Generally, medical X-ray detectors employing a scintillating phosphor screen to absorb X-rays and produce light suffer the loss of spatial resolution due to lateral light diffusion in the phosphor screen. To reduce lateral light diffusion and maintain acceptable spatial resolution, the phosphor screens must be made sufficiently thin.
The spatial resolution and X-ray detection ability of an imaging apparatus are often characterized by the modulation transfer function (MTF) and X-ray absorption efficiency, respectively. Thin phosphor screens produce better MTF at the expense of reduced X-ray absorption. Usually, the coating density and the thickness of the phosphor screen are used in the design tradeoff between spatial resolution and X-ray absorption efficiency.
For example, the Lanex Fine and the Lanex Fast Back screens are two typical commercial screens manufactured by Eastman Kodak Co. Both are made of Gd2O2S(Tb) phosphor. The Lanex Fast Back screen is thicker, absorbs X-rays more efficiently, but has lower resolution than the Lanex Fine screen. On the other hand, the Lanex Fine screen is thinner than the Lanex Fast Back screen, absorbs X-rays less efficiently, but has higher resolution. The coating density of the Lanex Fine and the Lanex Fast Back screens are 34 mg/cm2 and 133 mg/cm2, respectively. The Lanex Fine and the Lanex Fast Back screens have an X-ray absorption efficiency of 24% and 63% (for 80 kVp, with tungsten target, 2.5-mm Al inherent filtration, and filtered by 0.5-mm Cu) and an MTF value of 0.26 and 0.04 at 5 c/mm, respectively.
In order to improve X-ray absorption and maintain spatial resolution, the use of double screens in conjunction with a double-emulsion film has been incorporated in conventional screen-film (SF) radiographic apparatus. Similarly, the dual-screen technique has also been used in computed radiography (CR) to improve the X-ray absorption efficiency. In the digital CR apparatus, a storage phosphor screen is used in place of the prompt emitting phosphor screen employed in the SF apparatus. No film is needed. Upon X-ray exposure, the storage phosphor screen stores a latent image in the form of trapped charge that is subsequently read out, typically by a scanning laser beam, to produce a digital radiographic image.
Recently, digital flat panel X-ray imagers based upon active matrix thin film electronics have become a promising technology for applications such as diagnostic radiology and digital mammography. There are two types of X-ray energy conversion methods used in digital radiography (DR), namely, the direct and indirect method. In the direct method, the X-rays absorbed in a photoconductor are directly transduced into a charge signal, stored on the pixel electrodes on an active matrix array (AMA) and read out using thin film transistors (TFTs) to produce a digital image. Amorphous selenium (a-Se) is usually used as the photoconductor. In the indirect method, a single phosphor screen is used to absorb X-rays and the resultant light photons are detected by an AMA with a single photodiode (PD) and a TFT switch at each pixel. The photodiode absorbs the light given off by the phosphor in proportion to the X-ray energy absorbed. The stored charge is then read out, like the direct method, using the TFT switch.
Hydrogenated amorphous silicon (a-Si:H) is commonly used to form the photodiode and the TFT switch. FIG. 1A shows a cross-section (not to scale) of single imaging pixel 10 in a prior art a-Si based flat panel imager. Each imaging pixel 10 has, as shown in FIG. 1B, photodiode 70 and TFT switch 71. A layer of X-ray converter (e.g., luminescent phosphor screen 12) is coupled to the photodiode-TFT array. Photodiode 70 comprises the following layers: passivation layer 14, indium tin oxide layer 16, p-doped a-Si layer 18, intrinsic a-Si:H layer 20, n-doped a-Si layer 22, metal layer 24, dielectric layer 26, and glass substrate 28. X-ray photon path 30 and visible light photon path 32 are also shown in FIG. 1A. When a single X-ray is absorbed by the phosphor, a large number of light photons are emitted isotropically. Only a fraction of the emitted light reaches the photodiode and gets detected.
FIG. 1B shows a block diagram of the flat panel imager 80, which consists of a sensor array 81. The a-Si based sensor array consists of m data lines 84 and n gate lines 83. Each pixel consists of a-Si photodiode 70 connected to thin film transistor (TFT) 71. Each photodiode 70 is connected to common bias line 85 and the drain (D) of its associated TFT. Bias lines 85 carry bias voltages applied to photodiodes 70 and TFTs 71. TFTs 71 are controlled by their associated gate lines 83 and when addressed, transfer stored charge onto data lines 84. During readout, a gate line is turned on for a finite time (approximately 10 to 100 μs), allowing sufficient time for TFTs 71 on that row to transfer their pixel charges to all the m data lines. Data lines 84 are connected to charge amplifiers 86, which operate in parallel. In general, charge amplifiers 86 are divided into a number of groups, with each group typically having 32, 64, or 128 charge amplifiers. The associated charge amplifiers in each group detect the image signals, and clock the signals onto multiplexer 87, whence they are multiplexed and subsequently digitized by an analog to digital converter 88. The digital image data are then transferred over a coupling to memory 93. Gate lines 83 are turned on in sequence, requiring approximately a few seconds for an entire frame to be scanned. Additional image correction and image processing are performed by computer 90, and the resulting image is displayed on monitor 91 or is printed onto media by printer 92.
To reduce electronic noise as much as possible, a correlated double sampling (CDS) circuitry 89 may be disposed between each charge amplifier 86 and multiplexer 87. In the readout sequence of image signals, the charge signal on each data line 84 is sampled before and after the detection of signal charge by charge amplifier 86, and the resulting difference becomes the signal measured. In this sampling scheme, the background noise is reduced from the image signal. Double correlated sampling circuits 89 are preferable for flat panel imager 80, but are not needed for the imager to function. The synchronous operations of the various units of flat panel imager 80, namely, gate drivers 82, charge amplifiers 86, correlated double sampling circuits 89, and analog-to-digital converters 88, which provide the necessary timing, biasing, switching, sampling, scanning, and data readout functions, are controlled by computer 90 via control logic unit 94.
The operation of the a-Si based indirect flat panel imager is known by those skilled in the art, and thus only a brief description is given here. Incident X-ray photons are converted to optical photons in the phosphor screen 12, and these optical photons are subsequently converted to electron-hole pairs within the a-Si:H n-i-p photodiodes 70. In general, a reverse bias voltage is applied to the bias lines 85 to create an electric field (and hence a depletion region) across the photodiodes and enhance charge collection efficiency. The pixel charge capacity of the photodiodes is determined by the product of the bias voltage and the photodiode capacitance. The image signal is integrated by the photodiodes while the associated TFTs 71 are held in a non-conducting (“off”) state. This is accomplished by maintaining the gate lines 83 at a negative voltage. The array is read out by sequentially switching rows of TFTs to a conducting state by means of TFT gate control circuitry. When a row of pixels is switched to a conducting (“on”) state by applying a positive voltage to the corresponding gate line 83, charge from those pixels is transferred along data lines 84 and integrated by external charge-sensitive amplifiers 86. The row is then switched back to a non-conducting state, and the process is repeated for each row until the entire array has been read out. The signal outputs from external charge-sensitive amplifiers 86 are transferred to analog-to-digital converter (ADC) 88 by parallel-to-serial multiplexer 87, subsequently yielding a digital image. Alternatively, individual ADCs can be located at each signal output from charge amplifier 86. Multiplexer 87 could thus be removed from flat panel imager 80. The flat panel imager is capable of both single-shot (radiographic) and continuous (fluoroscopic) image acquisition.
Another imaging technique, known as dual energy subtraction imaging, has been used to reduce the impact of anatomic background on disease detection in chest radiography and angiography. This method is based on the different energy-dependent absorption characteristics of bone and soft tissue. In general, two raw images are produced. One is a low-energy and high-contrast image, and the other is a high-energy and low-contrast image. By taking nonlinear combinations of these two images, pure bone and soft-tissue images can be obtained. This imaging technique would improve diagnosis of pathology and delineation of anatomy.
The dual energy subtraction imaging has two general approaches: dual-exposure technique and single-exposure technique. In the dual-exposure technique, two different images are obtained from a detector by making two exposures at two different X-ray tube voltage settings. Since a double exposure of the patient must be performed, and the switching of the X-ray tube voltage must take a finite time, the double exposure technique would be sensitive to patient motion artifacts and misregistration between the two images. In the single-exposure technique, in which an energy filter is sandwiched between two detectors to attenuate the low-energy component, two different images are simultaneously obtained by making only one exposure of the patient. The single-exposure technique has the advantages of reducing patient motion misregistration artifacts and reducing X-ray dosage. The dual energy subtraction imaging has been implemented in both the screen-film and computed radiography apparatus with either the single-exposure or the dual-exposure technique.
Prior art screen-film apparatus 40 as shown in FIG. 2 combines asymmetric screens (front screen 44 and back screen 56) with a zero-crossover film coated with asymmetric emulsions (i.e., there is no light emitted by each screen crossing the film support to expose the emulsion on the other side) as shown in FIG. 2. X-ray photon path 42 illustrates the incoming path of X-rays to screen-film apparatus 40. Anti-crossover layers 48 and 52 (light absorbing layer) are deposited between each emulsion layer (front emulsion layer 46 and back emulsion layer 54) and film support layer 50. Relatively slow, high-resolution front screen 44 exposes high-contrast front emulsion layer 46. The combination of front screen 44 and front emulsion 46 is primarily responsible for imaging the lung fields. In addition, fast, standard-resolution back screen 56, which absorbs X-ray quanta with high efficiency, exposes wide-latitude back emulsion 54, and is primarily responsible for imaging the low-exposure mediastinum and retrocardiac regions. As a result, both the lung field and the mediastinum areas are clearly recorded. This screen-film imaging apparatus is primarily an analog (not digital) apparatus in which the exposed film must be chemically processed to form the final image. It can take a few minutes from exposure to image display. The apparatus has a narrow dynamic range and thus a narrow exposure latitude. The image cannot be digitally processed for image enhancement, displayed on monitor, stored in computer or digital storage devices, and transmitted wirelessly or via the Internet or other communications network.
Turning to FIG. 3, prior art dual-screen computed radiography (CR) imaging apparatus 60 has been used to improve the X-ray absorption and thus the overall image signal-to-noise ratio (SNR) and the contrast-to-noise ratio (CNR). With this technique, two CR screens (front screen 62 and back screen 64) are placed in a cassette for exposure by X-rays 66 through patient 68, as shown in FIG. 3. Exposed screens 62 and 64 are separately scanned with a laser scanner to form a front image and a back image which are then superimposed in various ways to optimize the quality of the final image.
For example, in chest imaging one can use a high-resolution (HR) screen and a standard-resolution (ST) screen to improve the MTF without compromising on the total X-ray absorption efficiency, as compared to the use of a ST screen alone. The use of a ST screen in the back of a HR screen does not degrade the high frequency performance of the HR screen. Instead, it enhances the image quality in the low to medium frequency range and preserve the quality in the high frequency range. As a result, this technique allows high-resolution details to be imaged (e.g., pneumothorax or rib fractures).
One prior art multi-screen CR apparatus has a plurality of stimulable storage phosphor plates that are exposed to X-rays to record a radiographic latent image of a subject viewed from the same direction. The image signals read out from the stimulable phosphor plates are superimposed to obtain an averaged image signal and to reduce the various noises associated with each component of the CR imaging apparatus. The averaged image signal is then subjected to a gradation process for enhancing the contrast of the image. As a result, the diagnostic efficiency and accuracy can be markedly improved.
Another prior art dual-screen CR apparatus for producing X-ray images of a subject exposed to an X-ray beam has a storage phosphor plate used for receiving X-ray radiation after passage through the subject. The storage phosphor plate has a substrate having two major faces with a photo-stimulable storage phosphor layer disposed on each face. The two storage layers have different materials or thicknesses. A double readout device is used to read out the radiation signals stored in each storage layer. The double readout device includes two separate sets of readouts apparatus for reading out each of storage layers. Each readout apparatus has a scanning laser beam to excite the storage charges, an optical collector to collect the stimulated light, a photomultiplier tube to convert the stimulated light into an electrical signal. The substrate (such as a metal foil) is impenetrable by the exciting laser beams and the stimulated light so that there will be no crossover of image signals generated in the two storage layers.
Other prior art dual-screen CR apparatus utilize an image superposition method wherein an addition process is carried out on a plurality of image signals representing the two radiation images of a single object recorded by two storage phosphor screens. The image signals corresponding to the picture elements are weighted with predetermined weight factors and then added to form the output image signal. The value of the weight factor with respect to the frequency components, which have a low signal-to-noise ratio, is rendered smaller than the value of the weight factor with respect to the frequency components, which have a higher signal-to-noise ratio, in accordance with the frequency characteristics of each of the image signals.
As a digital imaging technology, the above-mentioned CR-based imaging apparatus possess the desirable digital advantages over the screen-film apparatus. However, the CR apparatus needs a laser scanner to convert the latent x-ray image into the output digital image, and an optical unit to erase the residual image left on the storage phosphor plate from the previous X-ray exposure. This can take about 30 seconds to a few minutes for an image to be displayed.
Other prior art apparatus relate to dual energy subtraction imaging. One prior art dual energy subtraction technique uses a conventional screen-film combination to detect calcification in solitary pulmonary nodules. In this technique, two exposures are used. However, in clinical practice patient motion during the interval between the two radiographic exposures will degrade the subtraction image. Another prior art technique implements a single-exposure technique in dual-energy chest radiography using one X-ray tube voltage and two different screen-film combinations, separated by a copper filtration and loaded into a single cassette. This technique has been used to detect lung nodules by suppressing the bone contrast in chest radiography.
In the above-mentioned techniques using screen-film combinations, the image pair recorded on film has to be first digitized and then processed to produce the bone-free and soft tissue-free images. This would inherently reduce the throughout of the dual-energy radiographic procedure and reduce the image quality due to the image degradations caused by the film digitizer.
One prior art computed radiography system relies on a single-exposure dual energy subtraction imaging technique using scanning laser-stimulated luminescence. This apparatus produces the low-energy image and high-energy image simultaneously. These two original images are the images recorded by the first imaging plate (closer to the patient) and the second imaging plate, respectively. The new image is obtained by the subtraction processing. However, image magnification factor is slightly different between the first imaging plate and the second plate due to the fixed separation between the two plates. Therefore, in areas distant from the center of the X-ray beam, misregistration occurs.
Another prior art dual-exposure dual energy subtraction imaging apparatus with computed radiography uses a filter changer and an imaging plate changer to record a low-energy image and a high-energy image.
Although the above-mentioned dual energy CR apparatus have provided new diagnostic information that is not furnished by the conventional screen-film apparatus, these apparatus have been hampered by poor subtraction effectiveness, workflow inconveniences, and Detective Quantum Efficiency (DQE) limitations of the CR technology.
Recently, the dual-exposure dual energy subtraction imaging has also been implemented in digital flat-panel imaging apparatus based on a CsI:Tl scintillator coupled to an amorphous silicon TFT array or a Gd2O2S scintillator coupled to four CCD cameras. Although these dual-exposure based DR apparatus have shown improvements in the detection and characterization of lung diseases, there still are issues such as X-ray tube loading, X-ray dosage on patient, and patient motion artifacts.
As such, there exists a need for extending the application of dual scintillating screens (scintillating phosphor layers) to an indirect digital radiography (DR) apparatus. Moreover, there exists a need for extending the application of dual scintillating screens in an indirect digital radiography (DR) apparatus for the single-exposure dual energy subtraction imaging.